1. Field of the Invention
The present invention relates in general to magnetic resonance tomography (MRT), as used in the field of medicine for the examination of patients. The present invention relates in particular to a user interface for the correct planning or positioning of slice packets in the spatial domain based on an already-produced, corrected (equalized) MRT overview image.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear spin resonance, and has been used successfully as an imaging method in medicine and in biophysics for more than 15 years. In this examination method, the subject is exposed to a strong, constant magnetic field. As a result, the nuclear spins of the atoms in the subject, which were previously oriented in a random manner, come into alignment. Radio-frequency energy can now excite these “ordered” nuclear spins to a particular resonance. This resonance produces the actual measurement signal in the MRT, which is received by means of suitable receive coils. By the use of non-homogenous magnetic fields, generated by gradient coils, the signals from the examination subject can be spatially coded in all three spatial directions; in general, this is known as “spatial coding.”
The reception of the data in MRT takes place in k-space (domain frequency). The MRT image in the image domain, as it is called, is linked with the MRT data in k-space by means of a Fourier transformation. The spatial coding of the subject, which spans (fills) k-space, is accomplished by magnetic gradients in all three spatial directions. Distinctions are made between the slice selection (which defines a slice to be recorded in the subject, conventionally the z-axis), the frequency coding (defines a direction in the slice, conventionally the x-axis), and the phase coding (determines the second dimension within the slice, conventionally the y-axis).
Thus, first a slice is selectively excited, for example in the z-direction. The coding of the spatial information in the slice takes place by a combined phase and frequency coding by means of these two already-mentioned orthogonal gradient fields, which, in the example of a slice excited in the z-direction, are produced in the x-direction and y-direction by the aforementioned gradient coils.
A first possible sequence for recording the data in an MRT experiment is shown in FIGS. 2A and 2B. The sequence is a spin echo sequence. In this sequence, the magnetization of the spins is flipped into the x-y plane by a 90° excitation pulse. Over time (½ TE; TE is the echo time), there occurs a dephasing of the magnetization components, which together form the transverse-magnetization in the x-y plane Mxy. After a certain time has passed (e.g., ½ TE), a 180° pulse is emitted in the x-y plane in such a way that the dephased magnetization components are mirrored without modifying the direction of precession and speed of precession of the individual magnetization components. After a further time span ½ TE, the magnetization components again point in the same direction, i,e., there is a regeneration of the cross-transverse, designated a “rephasing.” The complete regeneration of the cross-transverse is designated the spin echo.
In order to measure a complete slice of the subject to be examined, the imaging sequence is repeated N times for different values of the phase coding gradient e.g. Gy, and in each iteration of the sequence the frequency of the magnetic resonance signal (spin echo signal) is sampled, digitized, and stored by the Δt-clocked ADC (analog-digital converter) N times at equidistant time intervals Δt, in the presence of the read-out gradient Gx. In this way, a numerical matrix as shown in FIG. 2B is obtained, produced row-by-row and having N×N data points. A symmetrical matrix having N×N points is only one example; asymmetrical matrices, or other k-space occupations, can also be produced. From such data sets in k-space, Fourier transformation is used to immediately reconstruct MR images of the relevant slice, having a resolution of N×N pixels,
The readout must be concluded in a time interval that corresponds to the decay of the transverse-magnetization. Otherwise, for example in a single-shot EPI sequence, the various rows of the k-matrix would be differently weighted in a manner dependent on the sequence in which they were acquired, certain spatial frequencies would be overemphasized, while others would be underemphasized. High measurement speeds, however, place extremely high demands on the gradient system. In practice, gradient amplitudes of approximately 25 mT/m are used. In particular for the change of polarity of the gradient field, significant energies must be converted in a very short span of time; the switching times are, for example, approximately 0.3 ms. The time in which the maximum gradient amplitude can be reached is known in general as the slew rate. The slew rate is, practically speaking, the speed with which a gradient field can be activated. Older systems have, or had, slew rates of 20-40 mT/ms. Modem systems have slew rates of 100-200 mT/ms, with the result that in modern systems, due to the gradient coil inductance, the respective gradient field has strong non-linearities.
In general, non-linearities of the gradient fields cause a distortion of the reconstructed MRT images, which undesirable in most cases. In modem MRT systems, such distortions can be corrected in the display of the image. The correction primarily serves cosmetic purposes, and does not increase the precision of the diagnostic findings. Most MR technicians, however, prefer to implement or activate this distortion correction, in particular if the obtained MRT images are to be forwarded to other specialist physicians, who may not be familiar with the details of MRT because such persons otherwise may consider the non-corrected images to be sub-par.
If the operator/technician now wishes to plan further measurements based on such a corrected image by the positioning of additional slice packets, a conflict arises because the planned slice packets, at their planned spatial position, do not actually “see” non-linear gradient fields, and thus the image plane that is planned based on the corrected image does not correspond to the actually recorded image plane.
Non-linearity and slew rate are directly connected with one another. Thus, for example there are MRT systems with a gradient system that can easily (for user-related reasons) be operated in two states (modes). The gradient system is constructed such that in a first operating mode (mode 1), a large but not strong gradient field can be produced with a relatively slow, moderate slew rate. Such a gradient field is as a rule very linear. In a second operating state (mode 2), however, a relatively small but strong gradient field can be produced with a rapid slew rate. As a rule, the gradient field produced in this way strongly non-linear.
If in mode 1 a first slice packet is recorded as an overview image, on the basis of which data acquisition (scans for scanned) further slices are then planned, but which are to be recorded in mode 2, a conflict as set forth above again results. The planned slice planes will not agree with the already-recorded image plane, due to the different non-linearity of the gradient fields.